A variety of implantable medical devices have been developed to treat diseased, injured, or deformed body conduits. In cases where malfunctioning body conduits have reduced inner diameters, there is usually reduced flow of vital fluid or gas through the conduits. In extreme cases, the malfunctioning conduits are occluded. Implantable medical devices used to open and/or expand, or to otherwise treat obstructed or constricted body conduits often reside in the conduits for a period of time following deployment of the devices. The devices serve primarily as mechanical supports inside the malfunctioning conduits to help maintain the conduits in more open, or patent, conditions.
These types of implantable medical devices are usually threaded through a healthy body conduit with a catheter, or other delivery mechanism, to a diseased area of the body conduit where the devices are employed. Many of these devices are frameworks made of a deformable metal suitable for implantation, such as stainless steel, cobalt-chromium, and other precious metals and/or alloys thereof. The devices are often employed with the aid of a balloon placed within the device framework that expands the framework until it presses against and engages the inside, or lumenal, wall of the body conduit. Other devices have combinations of configurations and material compositions that enable the devices to “self-expand” from a compacted shape to an expanded conformation without the aid of an inflatable balloon or other mechanical expansion means. Some devices are self-expanding under the influence of thermal or stored elastic energy alone. These devices are typically made with a shape-memory metal alloy commonly referred to as a super-elastic nickel-titanium composition (Nitinol). Shape-memory devices are intended to reduce or eliminate the need for an inflatable balloon.
Despite these and other advances in such implantable metal devices, the devices remain in the body conduit after palliation, treatment, or cure has been effected. While leaving the device in place may be benign in some situations, in other situations it would be preferred if the device did not remain in the body conduit. Of particular concern in leaving a metallic implant in a patient is the negative impact a subsequent procedure may have on the patient as a result of the implanted device. For example, the strong magnetic field produced by a magnetic resonance imaging (MRI) machine can adversely interact with the metallic implant. Also, the presence of a chronically implanted device can cause abrasions or erosions to tissue in which the device is implanted. Additionally, a subsequent endoluminal procedure can also result in an adverse encounter, or interaction, with a previously implanted device.
One approach to removing an implantable medical device from a body conduit has been to construct the device from certain polymeric materials that are either absorbed or degraded by physiological processes of the implant recipient. Bioabsorbable implants are constructed of polymeric materials designed to be benignly absorbed by the body over time. Materials capable of bioabsorption are also referred to as “absorbable,” “bioresorbable,” “resorbable,” “degradable,” and “biodegradable.” All these terms are all considered herein as synonymous with the term bioabsorbable with regard to the present invention.
The most common, bioabsorbable polymeric materials are removed from an implant recipient by hydrolysis of the polymeric material into metabolites, or break down products, that are substantially non-toxic to the implant recipient. The absorption of bioabsorbable polymeric materials typically begins by exposing the bio-absorbable material to aqueous fluids and/or certain enzymes under normal physiological conditions. The bioabsorption process usually continues until the device is entirely gone from the body conduit, or other implant site.
The particular polymers used to make an implantable medical device determine many of the properties of the device. Of particular significance are the biocompatibility of the bioabsorbable polymer and breakdown products thereof, bioabsorption rate, mechanical compatibility and compliance with tissue in which the device is implanted, rate of expansion, if any, mechanical strength of the device, and geometrical design.
An example of a mechanically expandable degradable stent that is tissue compliant, but of poor strength is disclosed by Beck et al. in U.S. Pat. No. 5,147,385. The stent of Beck et al. is made of poly(ε-caprolactone) polymer. According to Beck et al., the poly(ε-caprolactone) polymer melts in a temperature range between forty-five (45) and seventy-five (75) degrees centigrade. This is said to confer an ability to melt and mold the stent to the body lumen as the stent is being deployed in the lumen with the assistance of a heated balloon catheter. Unfortunately, when the polymeric material used to make the stent is placed inside a body conduit and raised above its melt temperature to undergo a re-modeling process, any inconsistency or inadequacy in thermal transfer from the heated balloon to the device as it is deployed can cause irregular and potentially unpredictable device deformations. Consequently, implantable medical devices intended for use as degradable stents or other temporary scaffoldings for a body conduit made of poly(ε-caprolactone), or similar homopolymers that require a thermal softening or melting transition above normal human body temperature in order to expand, are likely to lack sufficient reliable mechanical strength to be practical devices.
Another problematic expandable biodegradable stent is disclosed by Healy et al. in U.S. Pat. No. 5,670,161. The Healy et al. stent is made of a copolymer of l-lactide and ε-caprolactone that needs to be heated near or above its glass transition temperature (Tg) to be expanded. The stent's in vivo mechanical strength is dependent on the properties imparted by the glassy state existing below the copolymer's Tg. The stent is described as expandable using a “thermally-assisted mechanical expansion process at a temperature between about 38 degrees centigrade and 55 degrees centigrade.” Hence, the Healy et al. stent does not expand at normal human body temperature of 37 degrees centigrade and is also described as risking a “potentially hazardous” fracture if expansion is attempted in the brittle and glassy state found below the Tg of the stent.
As with Beck et al., the Healy et al. stent must be reliably heated above normal human body temperature and undergo a thermal transition before it can be safely remodeled by plastic deformation, expand, and become deployed. Healy et al. indicate a stent can be fashioned from a copolymer of l-lactide and ε-caprolactone that achieves a balance between sufficient mechanical strength to support a body lumen and the ability to remodel and expand the stent just above normal human body temperature. While this may be the case, an expandable endolumenal device made of a copolymer of l-lactide and ε-caprolactone requiring a thermal transition above normal human body temperature will always require application of heat to the device in excess of normal body temperature while it is inside the body conduit.
A biodegradable lactic acid-based polymeric material exhibiting “shape memory” properties for use in constructing implantable medical devices is disclosed by Shikinami in U.S. Pat. No. 6,281,262. Among the various medical devices that can be made with the Shikinami material, expandable supports for body conduits are disclosed. The supports are initially made in a cylindrical shape and subsequently collapsed into a secondary shape at a raised temperature. Upon cooling below the glass transition temperature (Tg) of the polymer, the support remains in the secondary collapsed state. When the support is reheated to a temperature (Tf) higher than the glass transition temperature (Tg), but below the crystallization temperature (Tc) of the polymeric material, the support reverts to its initial shape.
The primary polymer is poly-d,l-lactic acid. The ratios of the lactic acid isomers used in the polymer can be varied to achieve different thermal transition properties of the final polymeric material. The primary polymer can be mixed, or blended, with other biodegradable or bioabsorbable polymers, such as crystalline poly-l-lactic acid, poly-d-lactic acid, polyglycolic acid, amorphous polydioxanone, polycaprolactone, or polytrimethylene carbonate. Shikinami points out that regardless of the particular polymeric material chosen for use in his invention, the “shape-recovering treatment” should be performed at a temperature above normal human body temperature between 45 degrees and 100 degrees centigrade.
Devices requiring in situ application of heat above normal human body temperature are problematical at best. In addition to the possibility of causing patient discomfort in the form of pain with such a procedure, local trauma to same tissue intended to be medically treated with the device is also a possibility. Furthermore, it is unknown whether such trauma to “lumenal tissue” will stimulate undesirable tissue processes to begin at the implant site, such as a hyperplasia.
Stinson in U.S. Pat. No. 6,245,103 discloses a self-expanding stent constructed from bioabsorbable filaments that is self-expanding without the application of heat. The stent is said to have radial strength similar to metal stents by virtue of the braided construction of the stent and the chemical structure of certain filaments. The tubular self-expanding stent of Stinson is made by helically winding and interweaving resilient filaments of a bioabsorbable material into a particular braided configuration. Radial strength is imparted to the braided stent through use of two sets of interwoven filaments acting upon one another to “create an outwardly directed radial force sufficient to implant the stent in a body vessel upon deployment from a delivery device.”
As the stent changes shape from a compacted to an expanded configuration, the radius of the stent is increased and the axial length is decreased. According to Stinson, the shortening of the device can be predicted and compensated for. While this may be the case, shortening of the stent may not be acceptable in some applications, such as cardiovascular applications.
A similar bioresorbable self-expanding stent made of braided filaments is disclosed by Jadhav in U.S. Pat. No. 6,368,346. The principle teaching of Jadhav is the use of blends of bioresorbable polymers instead of co-polymers to make the filaments. Blended polymers are preferred by Jadhav because the chemical composition of the filament material is said to be more easily adjusted with blending than by synthesizing a new batch of co-polymer material. Blending is also said to reduce batch to batch variation common with synthesized co-polymers. As with Stinson, supra, the braided filaments form a tubular structure that shortens in length upon radial expansion of the compressed stent.
Jadhav also discloses an embodiment of a radially self-expanding extruded tubular stent made with blended bioresorbable polymers. The stent has walls that may be populated with openings. As with the braided stent, Jadhav's extruded stent is also “axially retractable” and shortens in length as it increases in diameter and radially expands. Excessive shortening (e.g., greater than about ten percent (10%) during deployment can result in deployment inaccuracy and cause intragenic trauma to wall tissue of the body conduit.
Igaki discloses a stent in U.S. Pat. No. 6,200,335 that is intended to be more compliant with tissue in which the stent is implanted. In this pursuit, the walls of the stent are gradually decreased in thickness toward each end of the stent. Additional flexibility can be imparted to the stent by introducing holes in the tapered wall material. The combination results in a stent structure with a Young's modulus equal to or slightly larger than 3×107 pascal, which Igaki describes as approximately representative of a blood vessel. Though primarily directed to a metallic stent, Igaki indicates his tubular stent can be made by extruding or injection molding a polymer material that has “biological absorptivity.” None of Igaki's stents are said to be compressible or expandable. Rather, the stents are referred to simply as inserts. The advantages of an endoluminal device that is both compressible for transport through a body conduit and expandable for delivery of the device to an implantation site are well known and appreciated in the surgical arts.
Virtually all of the above-summarized bioabsorbable stents rely on the readily available α-hydroxy ester polymers and copolymers derived from glycolide and lactide. While these monomeric components provide a predictable resorption rate by virtue of their hydrolyzable bonds, the monomers contribute only limited amounts of freely rotating aliphatic component into the polymer chain. As a result, polyglycolic acid (PGA) possesses a relatively high glass transition temperature (Tg) of approximately 36° C. (Benicewicz B C, Hopper P K., “Polymer for Absorbable Surgical Sutures—Part I,” Journal of Bioactive and Compatible Polymers, 5:453-472 (1990)) while polylactic acid (PLA) delivers a Tg of approximately 52° C. (Middleton J C, Tipton A J., “Synthetic Biodegradable Polymers as Orthopedic Devices,” Biomaterials 21:2335-2346 (2000)). Additionally, both polyglycolide (PGA) and single isomeric forms of polylactide (d-PLA or l-PLA) carry the potential for crystallization, imparting both dimensional stability and increasing polymer rigidity.
To impart some improved flexibility to the polymeric chain, many of the above-summarized references utilize either blending or co-polymerization of the α-hydroxy ester with polycaprolactone, a polyester which utilizes a monomer possessing a larger aliphatic component of five successive methylene groups. This increased aliphatic component provides polycaprolactone with greater amorphous state flexibility and a significantly lower Tg than the α-hydroxy esters. While the rotational features of this relatively extended aliphatic chain component impart improved chain mobility and a resulting Tg of approximately minus sixty degrees centigrade (−60° C.) (Middleton J C, Tipton A J., “Synthetic Biodegradable Polymers as Orthopedic Devices,” Biomaterials 21:2335-2346 (2000)), this rotational benefit is somewhat counteracted by polycaprolactone's tendency to crystallize with the accompanying rigidity.
It is these combined structural properties of the α-hydroxyester and caprolactone-based polymers that impart the specific thermal and flexural characteristics to these bioabsorbable stent constructions that require application of heat above normal body temperature and/or extrinsically applied force to deploy the device. It is noteworthy that polycaprolaction homopolymer has particularly longevity in vivo.
In addition to the inherent mechanical weakness conferred on the above-summarized devices by the polymers and processes from which they are formed and deployed, as discussed above, there are dangers inherent in applying heat to implantable devices, particularly endoluminal devices, at the implantation site. It would be desirable, therefore, if the polymeric material of an implantable medical device could reliably change shape, or conformation, without the need for the material to undergo a thermal transition, and especially without need for application of heat to the device at the implantation site. A non-elongating, self-expanding, implantable medical device made of biocompatible materials allowing the device to assume an expanded configuration when deployed at or near body temperature has numerous clinical benefits.
There is a need, therefore, for a non-elongating self-expanding support for a body conduit made of materials configured to allow the support to be reliably deployed inside or outside a body conduit at, or below, normal human body temperature. There is a further need for a support for a body conduit that is bioabsorbed into the body within approximately one year of implantation to assure complete removal of the support from the implantation site and, if needed, provide the opportunity in a subsequent procedure to implant another device.
None of the above-summarized stents meet the needs of an implantable medical device for use in a body conduit that is constructed from a non-blended hydrolyzable polymeric material as an integral flexible fenestrated framework that does not substantially change axial length with radial compression and expansion of the stent and is self-expanding at, or below, normal human body temperature without the polymeric material undergoing a thermal transition. Nor do the devices provide for variable bioabsorption rates in different parts, segments, or portions of the devices.